Spectral domain optical coherence tomography system

ABSTRACT

An optical coherence tomography device is disclosed for improved imaging. Reduced levels of speckle in the images generated by the device are obtained by forming a B-scan from a plurality of A-scans, wherein each resolution cell of the B-scan is generated through compounding of a subset of the A-scans and wherein at least some of the subset of A-scans are separated by at least half the diameter of a speckle cell both tangent to and orthogonal to the B-scan at that cell.

PRIORITY

This application is a continuation of U.S. application Ser. No.13/312,245, filed Dec. 6, 2011, which is a divisional of U.S.application Ser. No. 12/842,935, filed Jul. 23, 2010 (now U.S. Pat. No.8,085,408), which is a divisional of U.S. application Ser. No.11/820,773, filed Jun. 20, 2007. This application claims the benefit ofthe filing date under 35 U.S.C. §119(e) of Provisional U.S. PatentApplication Ser. No. 60/815,107, filed on Jun. 20, 2006, and ProvisionalU.S. Patent Application Ser. No. 60/925,104, filed on Apr. 18, 2007,which are hereby incorporated by reference in their entirety.

TECHNICAL FIELD OF THE INVENTION

The subject invention relates to diagnostic and measurement devices forevaluating a patient's eye. In particular, a spectral domain opticalcoherence tomography system is disclosed.

BACKGROUND OF THE INVENTION

Optical Coherence Tomography (OCT) is a technology for performinghigh-resolution cross sectional imaging that can provide images oftissue structure on the micron scale in situ and in real time. OCT is amethod of interferometry that uses light containing a range of opticalfrequencies to determine the scattering profile of a sample. Opticalcoherence tomography (OCT) as a tool for evaluating biological materialswas first disclosed in the early 1990's (see U.S. Pat. No. 5,321,501 forfundus imaging.). Since that time, a number of manufacturers havereleased products based on this technology. For example, the assigneeherein markets a device called the StratusOCT. This device is used fordiagnostic imaging and provides direct cross sectional images of theretina for objective measurement and subjective clinical evaluation inthe detection of glaucoma and retinal diseases. The device can generateimages of macular thickness, the retinal nerve fiber layer, the opticdisc, the cornea, and other parts of the eye. This device is based on aversion of OCT known as time domain OCT.

In recent years, it has been demonstrated that frequency domain OCT hassignificant advantages in speed and signal to noise ratio as compared totime domain OCT (Leitgeb, R. A., et al., Optics Express 11:889-894; deBoer, J. F. et al., Optics Letters 28: 2067-2069; Choma, M. A., and M.V. Sarunic, Optics Express 11: 2183-2189).

In frequency domain OCT, a light source capable of emitting a range ofoptical frequencies excites an interferometer, the interferometercombines the light returned from a sample with a reference beam of lightfrom the same source, and the intensity of the combined light isrecorded as a function of optical frequency to form an interferencespectrum. A Fourier transform of the interference spectrum provides thereflectance distribution along the depth within the sample.

Several methods of Frequency domain OCT have been described in theliterature. In spectral-domain OCT (SD-OCT), also sometimes called“Spectral Radar” (Optics letters, Vol. 21, No. 14 (1996) 1087-1089), agrating or prism or other means is used to disperse the output of theinterferometer into its optical frequency components. The intensities ofthese separated components are measured using an array of opticaldetectors, each detector receiving an optical frequency or a fractionalrange of optical frequencies. The set of measurements from these opticaldetectors forms an interference spectrum (Smith, L. M. and C. C. Dobson,Applied Optics 28: 3339-3342), wherein the distance to a scatterer isdetermined by the wavelength dependent fringe spacing within the powerspectrum. SD-OCT has enabled the determination of distance andscattering intensity of multiple scatters lying along the illuminationaxis by analyzing a single the exposure of an array of optical detectorsso that no scanning in depth is necessary. Typically the light sourceemits a broad range of optical frequencies simultaneously.Alternatively, in swept-source OCT, the interference spectrum isrecorded by using a source with adjustable optical frequency, with theoptical frequency of the source swept through a range of opticalfrequencies, and recording the interfered light intensity as a functionof time during the sweep (U.S. Pat. No. 5,321,501).

The commercial OCT systems typically include some form of scanning minorconfiguration to scan the light beam across the eye in a planeperpendicular to the propagation axis of the beam. The most commoninterferometer configuration for OCT is the Michelson interferometer[FIG. 1a of U.S. Pat. No. 5,321,501]. Michelson interferometers returnsome reference arm light to the source, causing a conflict between thedesire to set the reference level for best performance of the detector,and to set the reference level low enough to be below theback-reflection tolerance. Some alternative interferometer topologiesallow the reference path to be completely in fiber, allowing simpleconstruction. If the reference path is completely in fiber then thesample path length can be varied instead (U.S. Pat. No. 5,321,501).

Non-reciprocal optical elements in the source arm [U.S. Pat. No.6,657,727 issued to Izatt, et al.] have been used to divert thereflected light that would otherwise return to the source to a detector.While this protects the source and increases its longevity,non-reciprocal optical elements in the source arm add significant coststo the interferometer manufacture.

Interferometers with topology different from the common Michelsontopology have been proposed for OCT (U.S. Pat. No. 5,321,501 FIG. 10,U.S. Pat. No. 6,201,608 issued to Mandella, et al., and U.S. Pat. No.6,992,776 issued to Feldchtein, et al.). Some of these designs route thereference light without retro-reflecting or otherwise reversing thereference light back toward the source. In such interferometer designssome light returned from the sample can reach the source, but this isless of a concern because in many applications only a small fraction(10⁻⁴ to 10⁻¹⁰) of the incident light is scattered from the sample andreturned to the interferometer.

There has been a continuing effort in the industry to improve theexisting OCT systems. For example, when measuring living tissue such asan eye, movement during the measurement period can cause a wide varietyof difficulties. Efforts have been made to increase the speed of datacollection to reduce the effects of motion of the subject. In addition,various approaches have been suggested to measure sample motion and thencompensate for that motion.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is an optical path system diagram including the OCT path, the LSOpath, the Fixation path, the Iris Viewer path, and the internal testtarget path.

FIG. 2 is a system block diagram including the OCT system, the LSOsystem, the Fixation system, the Iris Viewer system, and the motorizedchin rest.

FIGS. 3 a and 3 b are a two-part electrical block diagram of the PC andsystem peripherals.

FIG. 4 is a polarization paddle design.

FIG. 5 is a flow chart of scan states during the operation of theinstrument.

FIG. 6 is an example internal test target.

FIG. 7 illustrates analysis of the internal test target.

FIG. 8 illustrates one topology of an interferometer for OCT.

FIG. 9 illustrates another topology of an interferometer for OCT.

FIG. 10 a illustrates the delay line alignment.

FIG. 10 b illustrates a consequence of delay line misalignment.

FIG. 11 illustrates a mounting block for the corner cube rail.

FIG. 12 illustrates an embodiment for bending fiber as a method ofbirefringence compensation.

FIG. 13 illustrates a system with a large number of optical componentsworking off the left side of a retinal conjugate, with the ocular lensand eye on the right hand side of the retinal conjugate.

FIG. 14 a is a table of steps used for alignment of an OCT imagingsystem.

FIG. 14 b is a table of optional steps used for alignment of an OCTimaging system.

FIG. 15 a illustrates a grid of A-scans for acquiring a 3-D OCT volume.

FIG. 15 b illustrates a thick B-scan achieved by combining 2 B-scans.

FIG. 15 c illustrates a thick B-scan achieved by combining 3 horizontalB-scans.

FIG. 15 d illustrates a thick B-scan achieved by combining 3 verticalB-scans.

FIG. 16 a illustrates a wiggle pattern for acquiring speckle reducedB-scans.

FIG. 16 b illustrates how the A-scans of a traditional B-scan relate tospeckle at a given depth.

FIG. 16 c illustrates how the A-scans acquired for speckle reducedB-scans relate to speckle at a given depth.

FIG. 17 illustrates the Optical lay out of a test eye.

FIG. 18 illustrates the cross section of a retina.

FIG. 19 a illustrates the traditional optical alignment of two opticaldevices without chromatic aberration alignment.

FIG. 19 b illustrates the traditional optical alignment of two opticaldevices with chromatic aberration alignment.

FIG. 19 c illustrates the preferred optical alignment of two opticaldevices with chromatic aberration alignment.

FIG. 20 shows, in pictorial form, a conventional scanning optical systemand fundus camera.

FIG. 21 shows, in pictorial form, one embodiment of the currentinvention.

FIG. 22 shows, in pictorial form, another embodiment of the currentinvention.

FIG. 23 shows, in pictorial form, yet another embodiment of the currentinvention.

DISCLOSURE OF THE INVENTION

This document is intended to describe a new OCT system being developedby the assignee herein. The principals and applications of the inventionare set forth in part in the description which follows, and, in part,will be obvious to those skilled in the art from the descriptionprovided herein. Further advantages may be learned by practice of theinvention. The scope of the invention is defined by the claims, whichincludes known equivalents and unforeseeable equivalents at the time offiling of this application. This new system is a spectral domain opticalcoherence tomographer including a spectrometer. The system also includesa line scanning ophthalmoscope (LSO) and an iris viewing system. Certainaspects of the individual sub-systems are unique. In addition, thecombination of these sub-systems is also unique.

Some of the inventive concepts being employed in the subject OCT systemhave been described in earlier filed patent applications which will bereferenced herein all of which are incorporated by reference. Thisdisclosure is intended to describe the overall system.

FIG. 1 is a schematic of the principal optical components of the system10 specially designed to generate diagnostic images of an eye 30. Thesystem 10 includes four primary sub-systems, an optical coherenttomography (OCT) system 20, a line scanning ophthalmoscope (LSO) 40, afixation system 50, and an iris viewer 60.

The OCT system is a spectral domain system generally of the typedescribed in the above-cited articles. The OCT system 20 includes a lowcoherence light source 101 which in this case is a super luminescentdiode that has bandwidth of about 800 to 900 nm with a center wavelengthof 840 nm. One choice for this device is the SD371 manufactured bySuperlum Diodes in Moscow. The output from the SLD 101 is directed intoa fiber based interferometer 100. An input fiber 110 is connected to afirst port I of a 70/30 optical coupler 111. The coupler 34 directsthirty percent of the light out of port III to the sample arm fiber 112and seventy percent of the light out of port IV to the reference armfiber 113. The reference arm fiber 113 is optionally connected to anoptical attenuator 119. An optical attenuator is useful for attenuatingexcess signal passing through the reference arm. A variable opticalattenuator can also compensate for the variability of other components.In particular, a variable optical attenuator can be optimized duringsystem manufacture to account for attenuation differences between partsand ensure that the signal transmitted to port I of a 99/1 coupler 131is sufficient. This configuration defines a transmissive reference path,which has an advantage over Michelson interferometers in that thereference light is not returned back to the source 101.

A photodetector 132 in the interferometer monitors light coupled out ofcoupler 131. Alternatively, a monitoring detector could be positioned inthe reference arm fiber 113. The photodetector 132 is used to measurethe power of the source both for eye safety purposes and to monitor thedegradation of the source.

Sample arm fiber 112 directs light into a delay line implemented usingcorner cube 124. The corner cube is translatable along an axis asindicated by arrows A to change the path length of the sample arm. Thepath length of the sample arm is adjusted with respect to the referencearm to select the depth in the tissue at which the OCT image will becentered. The light exiting the corner cube 124 is directed to a pair ofscanning galvanometer minors 116 for scanning the beam in a planeperpendicular to the propagation axis of the beam. The light is thenpassed through a lens doublet defined by lens 150 and 152. A turningmirror 151 is interposed between the lenses. It is preferable that thespacing between lenses, and optical path length within lenses, along theOCT beam, such as between lenses 150 and 152, be greater than the freespace OCT depth range so that reflections off the surfaces of the lenseswill not create interference effects that might be interpreted as comingfrom structures in the eye.

The light beam is then turned towards the eye using a dichroic beamsplitter 160. The dichroic beamsplitter 160 functions to directs lightfrom the OCT path to the common optical path used by various subsystemsand redirects OCT return light back along the OCT path while redirectingother light backscattered from the eye along a different path. The lightis directed into the eye with a lens 162. Lens 162 is designed tocorrect for spherical aberrations in the eye. In order to compensate forrefractive error, we adjust position of lens 162 with respect to theremainder of the optics. Light entering the eye is reflected back fromvarious structures in the eye such as retinal layers. The reflectedlight travels back on the same path to input port III of coupler 111.Coupler 111 directs seventy percent of the light reflected from thesource and returned to port II to the combining coupler 131 while thirtypercent of the light returned from the source returns to the lightsource through port I. Alternatively, the 70/30 coupler 111 may be an80/20 coupler or a 90/10 coupler.

Coupler 131 functions to combine light returning from the sample arm andarriving from the reference arm to create interference effects. Amajority of this combined light is directed out of the coupler 131 atport III to a spectrometer 200. Further information about interferometerdesigns having a transmissive reference path can be found below in thesection APPARATUS FOR OPTICAL COHERENCE TOMOGRAPHY.

A polarization paddle is provided to optimize signal strength. OCTdepends on interference between sample and reference beams, and theinterference of these beams produces a modulation in power to the extentthat the polarizations of the beams match. Specifically, if one uses thecommon Poincaré sphere representation of polarizations, the amplitude ofthe interference fringes is proportional to the cosine of half the angleon the Poincaré sphere between the Poincaré sphere representations ofthe sample and reference polarizations.

Rotating birefringent elements are a common method of controllingpolarization. In fiber optics, bending the fiber is a convenient meansfor creating birefringence, and rotating the orientation of the bendsrotates the axis of birefringence. Such an assembly is often called apolarization paddle. (See, for example, chapter 9 of “Polarized Light inFiber Optics” Edward Collett, (c) PolaWave Group Lincroft N.J. 2003.)

Perfect polarization matching requires three polarization paddles tocompensate for arbitrary polarizations in the sample and reference arm.In practice, perfect polarization compensation is not required; onlysufficient polarization alignment is necessary to enable detection ofinterference. A single paddle is sufficient to compensate for mostpolarization differences seen in practice so that detection loss is nomore that 1-2 dB. The single paddle reduces equipment cost andsimplifies the design and control and improves exam throughputefficiency.

In the preferred embodiment, the single paddle is located in the samplearm but it could be located in the reference arm. A motor and hardwareis included for a rotatable paddle to provide polarization compensationfor the fiber. The fiber can be mounted with a U-shape bend onto thepaddle as shown in FIG. 4, or in the more traditional circular loop. Thepaddle can rotate out of the plane of the paper with the U-shaperemaining in one plane at all times. The design parameters are the threeradii (R₁, R₂ and R₃) and the three angles (α₁, α₂, α₃). Theseparameters are chosen to meet birefringence requirements as discussedbelow.

If the paddle is located in a single-pass portion of the interferometer,such as the reference path of a Mach-Zehnder design, then one expectstheoretically that a quarter-wave paddle will be effective in matchingpolarizations. Suppose we are given two arbitrary polarizations and anadjustable quarter-wave plate affecting the first polarization. Imaginethe locations A and B of the polarization states on the Poincaré sphereif we removed the quarter wave plate. The goal is to move A as close aspossible to B by rotating A on the sphere by 90° about an equatorialaxis of our choosing. Choose an axis x which puts both points in thehemisphere x>0. Sighting along the z axis, we can see both points on thesame hemisphere, and want to move A by either +90° or −90°, whicheverwill bring A closer to B. This choice corresponds to choosing to placeeither the fast or slow axis of the wave plate in the directioncorresponding to +x. It is always possible to move the first point sothat it is (1) in the same x>0 hemisphere as the second point, and (2)within ±90° azimuthally about the z axis. The resulting distance betweenA and B is always less than 90°. The <90° result is best understood byvisualization, but can also be proven using the law of cosines on thethree directions A, B and x. The resulting fringe amplitude, then, is atleast cosine 45° or 71% of what it would be with optimally matchedpolarization.

If the paddle is located substantially at the end of a bi-directionalpath, then two passes through a rotatable one-eighth wave paddle willhave the same benefits as derived above for the single pass through aquarter-wave paddle. If the paddle is located in a bi-directional pathbut located such that the light experiences significant uncontrolledbirefringence, such as by the sample, experimentation and simulationhave shown that three-eights of a wave of birefringence more robustlyrestores the interference fringe amplitude.

The spectrometer 200 is of the type disclosed in U.S. patent Ser. No.11/196,043, filed Aug. 3, 2005, (publication 2007/0030483) incorporatedherein by reference. Briefly, the spectrometer is in a folded Littrowconfiguration. Light enters the spectrometer and is directed to agrating 230. Grating 230 is preferably blazed for 840 nm, withapproximately 1200 lines/mm to give adequate spectral dispersion. Lightreflected from grating 230 is directed to a pixel camera 250. Thedispersion of the grating and the imaging lenses discussed below arechosen to spread wavelength from approximately 800 to 900 nm over thesensitive length of camera 250. A set of three lenses 240, 210 and 220is located between the grating and the camera. The light passes throughthese three lenses both on the path to the grating and on the returnpath to the camera. While each of the lenses contributes to focusing andcorrection, lens 220 is the primary lens for focusing and collectinglight from the grating. The grating is tilted in a way to induce conicaldiffraction which causes the returning light beam to be displaced awayfrom the fiber input, slightly out of the plane of the figure, andtowards the camera. The primary function of lens 210 is to correct foreffects of conical diffraction. Lens 240 functions primarily as a fieldflattening lens.

The SD OCT system is capable of generating two- or three-dimensionalimages of the retina in a manner known in the prior art. The subjectsystem has some additional capabilities that will be described below.However, one added feature is the ability to rapidly switch betweenimaging the retina to imaging of the cornea. This capability is providedby including an extra lens 180 which can be moved into the path of thelight in the sample arm to permit focusing of the light onto the cornea.In conjunction with the movement of the lens 180 into position, theposition of corner cube 124 is changed to allow the path lengthdifference between the sample arm and reference arm to correspond to theposition of the cornea. An advantage of this system is that theinformation about the cornea can be easily obtained without having toreposition the patient.

As an alternative to moving corner cube 124, mirrors can be moved toswitch in and out an extra fold in the optical path length, asillustrated for example in FIG. 5 of U.S. patent application Ser. No.11/243,665, filed Oct. 5, 2005, (publication 2007/0076217) incorporatedherein by reference. Other alternatives to moving corner cube 124 arethe rapid-scan optical delay (RSOD) devices disclosed in InternationalPatent Application No. WO 2005/033624 and in U.S. Pat. No. 6,654,127,either of which can be configured to provide a change in group-delay,with relatively small phase delay. The small phase delay is advantageoushere because changes in phase delay move the interference fringes acrossthe camera 250, which reduces the fringe intensity if the fringes moveduring exposure of camera 250. The group-delay devices with small phasedelay cause relatively less motion of the fringes the device is settlingafter a quick move.

While the optical delay does not need to be particularly rapid forcentering the imaging region, the ability to change the optical delayrapidly is useful for adjusting the optical delay on a scan-by-scanbasis. For example, an RSOD can be used to flatten the retina.Nominally, the retina will appear as a curved surface in a 3-D image ofthe eye. Using a predetermined delay profile, an RSOD can adjust theoptical delay on each A-scan and flatten the curved surface. Whensufficiently fast computation elements are available along with hardwarefeedback paths, on-the-fly optical delay adjustments can be computedfrom A-scan to A-scan. Alternatively, a sparse scan can obtain a selectcomplement of A-scans which can be used to identify the retina in eachA-scan and compute a fit to the retinal surface (say a spherical orparabolic fit), which can be used to generate a delay profile for theRSOD. Clearly, combinations of pre-determined profiles and on-the-flycomputations can also be used to direct the RSOD to modify the opticaldelay on an A-scan by A-scan basis in a SD OCT system.

The second main sub-system is a fundus viewer. The preferred fundusviewer technology is the line scanning ophthalmoscope LSO 40. The LSO 40includes a relatively narrow band light source which, in the preferredembodiment, is a super luminescent diode 410 emitting light at about 755nm with bandwidth about 5 nm. The light source is polarized. Light fromsource 410 is passed through shaping optics 415 to create a line oflight. The line of light is directed to a beam splitter 420 whichredirects the light to scanning galvanometer minor 430 for scanning theline in one axis perpendicular to the plane of propagation of the light.Beam splitter 420 comprises a reflective strip, so that illuminationfrom source 410, focused to a line along this strip, is directed to theeye, while light returning from the eye largely passes around this striptoward imaging lens 480. The illuminating light is directed to adichroic beam splitter 440 which is reflective of light at 755 nm andtransmissive at shorter wavelengths. The light is passed through a lensdoublet 450 to a dichroic beam splitter 460. Beam splitter 460 isreflective of light in the 700 nm wavelength region and transmissive forlight at 755 nm and 550 nm wavelengths. Beam splitter 460 needs to havehigh transmissivity at 755 nm only for the polarization of light used inthe LSO subsystem. The design of the dichroic coatings on beam splitter460 is easier if only one polarization state needs to be optimized.(Analogous design optimization is available for beam splitters 460 and160.) Light is then passed through beam splitter 160 into the eye. Beamsplitter 160 is reflective of light over 800 nm and transmissive oflight at shorter wavelengths. The beamsplitter 420 is nearly conjugateto the cornea, so that an image of the LSO light source reflected fromthe patient's cornea is formed on the reflective strip, thus blockingthis corneal reflection from the imaging optics.

Light from the LSO 40 is reflected by the eye and returns on the samepath to splitter 420. A portion of the reflected light is transmittedthrough splitter 420 and it is imaged via a lens 480 onto a line scancamera 490. Commercially available line scan cameras offering line ratesaround 10 kHz are appropriate for camera 490. As the galvanometer 430 isscanned, different portions of the retina are illuminated and imaged, sothat a two-dimensional image of the patient's retina is built up fromsuccessive exposures of the camera. With 512 lines in a frame, a framerate of 20 Hz is achieved. The scan range of galvanometer 430 is easilyvariable to adjust the field of view of the LSO.

The third main sub-system is an iris viewer 60. The iris viewer is usedprimarily to align the patient's eye with the optical axis of thedevice. The iris viewer includes an LED 610 positioned near lens 162 forilluminating the eye. Preferably, the LED generates light having awavelength of about 700 nm. The reflected 700 nm light is captured bylens 162 and travels back through splitter 160 to splitter 460 where itis reflected back through a series of lenses to a CMOS camera 620. TheLED can be polarized, or its output filtered by a polarizer, so that thelight reflected from the iris is largely polarized, and beam splitter460 optimized to reflect only one polarization state. Imaging the irisin polarized light has the side effect of revealing birefringence of thecornea. The camera 620 generates an output which is supplied to amonitor that will display an image of the iris. As discussed below, thisimage is used to position the patient.

The fourth main sub-system is a fixation system 50. Fixation system 50includes a display pad 510 for generating fiducial marks that will beprojected onto the patient's eye. The patient will be asked to fixateher eye on these fiducial marks. Pad 510 generates light at a visiblewavelength preferably between 450 and 600 nm. The light from pad 510 isconditioned by lens system 520 and directed through dichroic beamsplitters 440, 460 and 160 and focused into the eye via lens 162. Thepreferred fixation target is a variable sized, 2D fixation target. A 2Dfixation target provides both a center fixation target and the abilityto rapidly change visual stimuli for analysis of eye response.Preferably, the target size is variable from a point target to anoversized target embedded in a 120×120 pixel display covering a field ofview of 30 degrees.

As shown in the system block diagram of FIG. 2, a host computer 90 isused to interface between the operator 35 and the integrated system tocontrol the subsystems, either directly or through a controller such asa motor controller board 85, and to send and receive subsystem data,either directly or indirectly through and intermediary controller suchas a frame grabber board 45. As shown in the electrical block diagram ofFIG. 3, the host computer 90 provides the system with input devices(such as keyboard 91, mouse 92, or equivalents such as trackball orjoystick) and output devices (such as monitors 93 and printers 94) aswell as Input/Output devices such as digital storage devices such ashard drives (not shown), CDs (not shown), DVDs 95, etc., connectionports such as serial ports (not shown), parallel ports (not shown), USBports 96, fire-wire ports (not shown), and the like, and networkconnections 97 to local, peer-to-peer, distributed, or even theworld-wide web. The live iris camera 98 is preferably a direct PCperipheral, though it could also be integrated into the frame grabber,providing yet another image input to that device. The architecture ofthe frame grabber board can enable real time tracking by imageprocessing an image on the host PC and updating the galvo X and Yoffsets on the frame grabber in real time. In this embodiment, the galvoscan pattern is corrected and the appropriate region is imaged, even inthe presence of eye motion.

In order to improve the functionality of the device, a specific effortwas made to insure that the various sub-systems worked together in acooperative manner. For example, the OCT, LSO and iris viewer are alltelecentric systems, so that adjusting focus does not change themagnification of the image. The optical systems that focus on theretina, the LSO, OCT, and fixation, are parfocal so that they aresimultaneously in focus on the retina after compensation for refractiveerror. The systems that focus on the retina use different wavelengths,so their focus adjustments are calibrated to compensate for thedifferent focal lengths of the human eye at these various wavelengths.The systems that image the retina, OCT and LSO, are confocal systems,meaning that small areas of illumination are swept across the retina andimages of these areas directed to matched sensitive areas on thedetectors. Confocal imaging reduces glare from corneal reflections andscattering from other ocular media such as a cataract.

This design takes care to minimize polarization dependence in the opticsalong the OCT beam path. For example, differences in optical delaybetween the polarization states, known as polarization mode dispersion(PMD) cause OCT images with different depths for each polarizationstate. Given that the polarization state changes on transmission throughthe eye, polarization is difficult to fully control, and PMD generallyleads to broadening of the axial (depth) resolution in the OCT image.The dichroic beam splitters along the OCT path reflect, as opposed totransmit, the OCT beam because the beam reflected from dichroic coatingstypically has less PMD than the transmitted beam. Smaller polarizationdependent effects, such as fractions of a wave of birefringence, arealso controlled. The beam splitters are placed in locations where theOCT beam is telecentric, meaning the chief rays of the OCT beams forvarious positions of the scanner 116 are parallel, so that theangular-dependent polarization effects of the beam splitters do notchange as the OCT beam is scanned.

In use, the first step is to align the patient with the device. In thepreferred embodiment, the patient's head is put into a motorizedheadrest. A suitable headrest is described in U.S. patent applicationSer. No. 10/843,767, filed May 12, 2004 (publication No. 2005/0254009)which is incorporated herein by reference. The doctor will ask thepatient to view the fiducial marks generated by the pad 510. At the sametime, the doctor will observe the eye via a display (not shown)associated with the camera 620 of the iris viewer. Initially, thedistance between the patient's eye and the lens 162 is adjusted for bestfocus of the iris. Once the proper spacing has been achieved, theseparation between the lens 162 and eye is held constant while theposition of the eye with respect to the OCT system is varied to positionthe center of the OCT image at the desired depth within the eye. Lens162 is not carried by the motorized chin support 80 (FIG. 2). Rather, aseparate translation system is provided which is operatively linked tothe motion of the chin support during this positioning step. Furtherinformation about the approach used to position and align the patient'seye can be found below in the section “METHOD OF PATIENT ALIGNMENT FORMULTIFUNCTIONAL FUNDUS IMAGING”.

The primary purpose of the iris viewer is to help the operator centerthe patient's pupil so that the OCT and LSO beams pass through the pupilto the iris. A continuous view of the iris is helpful in keeping thepatient's pupil centered during retinal imaging. Note that the irisviewer can also be used to help position the OCT beam around cataracts.Further, it can be used to help collect OCT images through differentportions of the pupil, collecting light at different scattered anglesfrom the retina.

Once the patient has been aligned, a wide variety of OCT images can begenerated. The Fourier transform of the signals from the spectrometerprovide A-scan information at each X and Y position of the beam. (Insome methods scans are repeated at the same X and Y position to revealtime-dependent effects including pulsatile flow, Doppler shifts, etc.)This data can be collected and stored. Some A-scans can be acquired forpurposes other than imaging. For example, the scanning system 116 candirect the OCT beam in a circle outside the aperture of lens 162, duringwhich time the camera records the reference signal only, with no signalfrom the sample, thus collecting a background signal for use inprocessing. The processor can then generate and display other imageinformation (such as B-scans, en face images, Doppler images, etc.)familiar to the doctor. In addition to some of the more conventionalimaging modalities currently available on existing systems, the subjectapparatus has been configured to provide additional functionality.

For example, the system is configured to generate fundus type imagesbased on OCT data. This approach is described in U.S. patent applicationSer. No. 11/219,992, filed Sep. 6, 2005 (publication 2006/0119858) andincorporated herein by reference. In this type of analysis, theintensity information over the depth range for the OCT data at anyparticular X/Y location is integrated to generate a pixel in the fundusimage. The integration of intensity over a depth range to generate thefundus pixel may be performed by either accumulating intensities priorto compression (nominally logarithmic) for display or by compression ofintensities prior to accumulation. The fundus image can be continuouslydisplayed for the doctor to help interpret the OCT images and positionthe device. This fundus image based on OCT data is especially valuablefor registration of the location of the underlying OCT cross-sections,to an en-face view of the retina.

The OCT may be used to generate maps with three-dimensional rendering ofelevation, topographical maps or color or grayscale maps. U.S. patentapplication Ser. No. 11/717,263, Mar. 13, 2007, and incorporated byreference, discloses a variety of approaches including collectingcompound OCT scans for high definition scans and a data cube to providecontext for high definition scans. Also disclosed are standardizationtechniques for orientation, diagnostic metrics of texture andheterogeneity, retinal fluid maps, etc.

The section METHOD FOR COMBINING B-SCANS (“THICK B-SCAN”), below,discloses the concept of combining adjacent B-scans to reduce noise andspeckle and give an enhanced visual impression.

The software can be set up to generate elevation maps of tissue withrespect to fitted reference surfaces. This approach is described in U.S.patent application Ser. No. 11/223,549, filed Sep. 9, 2005 (publication2007/0103693) and incorporated herein by reference.

The system may also be set up so that the chromatic dispersion of thesample and reference paths are different from each other to create avariation in the relative group delay as a function of optical frequencybetween the sample and reference paths. Thereafter, the measuredinterference spectrum can be multiplied by a complex phase factor tocompensate for the mismatch. In this manner, the image contrast betweenreflections from the sample and image artifacts can be increased so thatthe doctor can better discern actual tissue images. Further informationon this approach is set forth in U.S. patent application Ser. No.11/334,964, filed Jan. 19, 2006, (publication 2006/0171503) incorporatedherein by reference.

As noted above, one problem associated with prior art systems relates toerrors resulting from the movement of the patient's eye during imaging.Errors of this type are reduced in the subject system because thescanning speed is much faster. For comparison, the time needed to scanthe eye using our current Stratus system is on the order of 2 seconds,while the subject system can cover the same scan region in only 0.026seconds.

The increase in speed is so great that new scanning sequences candirectly collect 3-D imaging data without the need for interveningtomograms. In some cases, performing scanning sequences collecting datain 2-D planar tomograms and then building a 3-D volume from the 2-Dslices is preferable because then existing software can be used forvisualization, reducing costs and time-to-market. Nonetheless, directcollection of 3-D voxel data in real-time using spectral domain opticalcoherence systems is now possible and the 3-D volume can be rendereddirectly for display.

In addition to increasing the scanning speed, other approaches have beendeveloped to still further reduce problems associated with eye movementduring measurement. For example, U.S. patent application Ser. No.11/331,567, filed Jan. 13, 2006 (publication No. 2006/0164653) andincorporated herein by reference discloses the concept of taking a fewpartial, fast OCT scans and using this information to provideregistration information during the slower, more complete OCT scans. Inanother approach, the LSO system 40 can be used to generate guidepoststhat can then be compared in the processor to the OCT images. Theinformation can be used to control the scanning of the galvanometerminors 116 in real time to compensate for patient eye movement.Alternatively, the LSO data can be used in post-processing to properlyregister the data acquired from the OCT system. More information aboutthis approach can be found in U.S. patent application Ser. No.11/389,351, filed Mar. 24, 2006 (publication 2006/0228011) andincorporated herein by reference.

It is also desired that the device exhibit long term repeatability andstability in the field. In the past, external targets where used by thedoctor to facilitate alignment and calibration. The subject system hasbeen provided with an internal calibration system to simplify theprocess of making sure the OCT and LSO systems are coaxially aligned.More specifically, a target 710 is provided which preferably includesfiducial marks such as crosshairs or horizontal and vertical alignmentbars (see FIG. 6). An auxiliary mirror 720 is provided located justbeyond the zone where the OCT and LSO beams are scanned when measuringthe patient. During a calibration step, the galvanometer mirrors 116 and430 are positioned so that the light from the OCT and LSO systems strikeminor 720 and are directed to target 710. The reflected light is imagedby the two respective detection systems. The driving systems for minors116 and 430 are adjusted until the images overlap. Information definingthe position of the galvanometer mirrors is stored and used to calibratethe device (see FIG. 7). The crosshairs in target 710 can be grooves tofacilitate imaging by the OCT device. Target 710 preferably includesreflectors at various depths for calibration of axial lengthmeasurements by OCT and confirmation of the axial resolution of the OCTsystem.

Yet another use for internal calibration is a galvanometer test. Afterscanning many times (up to billions of cycles), the motors of opticalscanning galvanometers mechanically wear. Before catastrophic failure,their bearings and lubricants deteriorate and affect performance. Thesemotors are typically driven by servo amplifiers that attempt to minimizethe difference between the actual galvanometer motor position and acommanded position. The actual position is provided either as an analogor digital signal. In ophthalmic scanning, image quality andrepeatability is directly dependent on galvanometer performance, so itis important to be able to characterize the closed loop response andadapt performance to achieve the desired response. The desired positionmay be achieved by adapting the loop filter of the servo or the commandsignal. In cases where there is no adequate internal alternativeavailable to achieve the required performance, the system can issue arequest service. The service request may be either through a notice tothe user on the system or a notice across a network to eitheradministrative personnel or directly to the service organization.

This internal calibration may be performed on an internal schedule, suchas monthly or weekly or on every n-th boot (where n is a positiveinteger), or through remote service and diagnostics.

The arrangement of the sub-systems leads to some novel combination. Forexample, and as noted above, the OCT system, LSO, iris viewer andfixation system are all parfocal. The iris viewer, which preferablydisplays a continuous image of the iris, greatly facilitates thealignment of the OCT system measurement system. This approach can becompared to the prior art approach, often used in fundus camera, ofusing the retinal imaging system, here an LSO, to first image the irisat a distance spaced significantly from the optimal positioningnecessary to obtain an OCT image. In order to then position the deviceto obtain a OCT and LSO images of the retina, the doctor would have tocarefully move the patient and imaging system closer together, along aline without deviation so that the imaging paths remain centered on thepupil. This adjustment would be difficult because the doctor would nolonger have the image of the iris displayed.

FIG. 2 is a block diagram of the overall system. It shows the operator35 interacts with the computer 90. The computer 90 acquires Iris Viewerimage data directly from the Iris Viewer 60 while fundus and OCT imageinformation are received through the frame grabber board 45.Alternatively, in some instances, one or more high performance graphicsboards can substitute for the frame grabber board 45 however, because ofthe multiple sources of image frames, specialized hardware waspreferable for the preferred design. The controller board 85 performsthe real-time system control for the OCT system 20, the fundus viewer40, the fixation target 50, the iris viewer 60, and the motorized chinrest 80.

FIG. 5 is a flow chart of one embodiment of the OCT scan states duringoperation of the instrument. The OCT system remains is a system idlestate until the operator indicates the start of a new acquisition 920.Since this chart is concerned with the scan states of the OCT system,the flow in this chart assumes that the iris image is already in focus.Then the first task is to align the scan. First we perform a backgroundscan 930 to determine current scan position. The scan is then aligned940 by setting the OCT reference depth using the Z-motor and moving thecombined ocular lens and eye with synchronized x-, y-, and z-motion(using the chinrest and Z-motor in combination) to optimize the focus ofthe fundus image. The system process 950 intermittently returns the OCTsystem to the background scan state 940 to ensure that the alignmentscan state 940 is performed using correct background information. Thesystem remains in the alignment scan state until the operator starts theacquisition state 970, unless returned to a background scan state or asystem timeout 960 occurs. The system timeout state 960 is entered aftera fixed interval in which the operator has not determined to acquiredata. In this embodiment, once the operator decides to acquire data,first a background scan is performed 980 and the acquisition scan isperformed 990. On acquisition scan completion, the system enters areview data state 995 and the operator can review the acquired datathrough various image display and analysis tools. On completion of thedata review, the system is returned to the system idle state 910.

Apparatus for Optical Coherence Tomography

The following embodiments describe interferometers for use in theinvention of record. Coherence-domain imaging techniques such as OCTpreferably use light sources with short axial coherence length, but withspatial coherence in the transverse directions. Superluminescent diodes,which are similar in structure to diode lasers, have short temporalcoherence and broad spatial coherence. By design, they do not laserbecause there is insufficient optical feedback. Superluminescent diodesare typically sensitive to optical back-reflection of output lightpotentially causing output power fluctuations and shortened lifetime.

The most common interferometer configuration for OCT is the Michelsoninterferometer. Most Michelson interferometers return some reference armlight to the source. The light returning to the source can be divertedby the use of non-reciprocal optical elements. To avoid the expense ofnon-reciprocal optical elements, one can control the polarization stateof the light and divert light returning to the source based on itspolarization state.

Some interferometer topologies allow the reference path to be completelyin fiber, allowing simple construction. Other interferometers usingessentially the same topology allow the reference path to be nearlycompletely in fiber, only deviating from continuous fiber to insertsimple free-space optics, such as a leakage optical attenuator. In OCT,the optical group delays must be approximately matched between sampleand reference paths. This matching is typically accomplished byadjusting the reference optical path length. If the reference path iscompletely in fiber then the sample path length can be varied instead,as noted in U.S. Pat. No. 5,321,501, c. 12, 11.16-21.

The OCT apparatus disclosed herein efficiently collects light from theeye, uses a reflective sample path, returns no reference light to thesource, and does not require circulators or other non-reciprocalelements.

Extreme split ratios in the fiber couplers can be avoided and oneconfiguration allows a safety monitor tap close to the sample arm tap.

FIG. 8 illustrates the topology of an interferometer for OCT thatreduces reflections of the light back into the source. Low coherencelight source 101 is typically a superluminescent diode (SLD) whichtypically tolerates back reflection of less that 10⁻³ of its outputlight. The SLD is coupled to source fiber 110 that routes light todirectional coupler 111 a. The optimal directional strength of thecoupling depends on system design choices and may be 90/10 (as shown inFIG. 8) or 70/30 (as shown in FIG. 1) or other as availability permits.Directional coupler 111 a splits the light into sample fiber 112 a andreference fiber 113 a. The sample path includes delay apparatus 114 toadjust the length of the sample path; shown in more detail in FIG. 9.The delay apparatus couples the light from fiber 112 a to a free-spaceOCT beam 115. Transverse scanner 116 deflects the OCT beam andpreferably creates a focus in the beam near the region of interest insample 30 a.

Some light scattered from sample 30 a returns through the scanner anddelay apparatus to sample fiber 112 a. Coupler 111 a routes this lightthrough loop 117 a to fiber coupler 131 a, where it is interfered withthe reference light. The combining coupler 131 a provides two outputs.These outputs could be used for balanced detection (U.S. Pat. No.5,321,501 FIG. 10) in which both detector 200 and detector 142 c areused to collect light for OCT. Alternatively, the coupling ratio ofcoupler 131 a can be adjusted to send most of the interfered light to asingle OCT detector 200. Each OCT detector can be a single photodetectorfor use in time-domain OCT or swept-source OCT, or a spectrometer foruse in spectral domain OCT.

Optional tap 121 diverts a fraction of the reference light to detector122, which may be used to monitor the source power. (Some reasons formonitoring include safety of the sample and detection of degradation inthe source 101.) The tap removes some fraction of optical power from thereference fiber 113 a, reducing the power that reaches coupler 131 a.Sensitivity in OCT can reach the shot-noise limit if the reference poweris large enough to bring the interference signal above receiver noise,but not so large as to bring intensity noise or beat noise above thelevel of shot noise. The reference power is approximately determined bythe source power, and the coupling ratios in directional couplers 111 aand 131 a, and adjusted by choice of tap 121.

The coupling ratios in directional couplers 111 a, 131 a and 121 arechosen to set a safe level of illumination to the sample, and to set theappropriate reference power at the detector or detectors. For example,in the case of ophthalmic OCT of the retina using light with wavelengthsnear 850 nm, the safe exposure level is approximately 0.5 mW, and theoptimum reference level at the detector is approximately 0.005 mW.Sources are available in this wavelength range having output power ofapproximately 5 mW. For these conditions one would use a coupling rationear 90%/10% in the splitting coupler 111 a so that 10% of the sourcepower reaches the sample. 90% of the scattered light will then be routedto loop 117 a. In the case where there is a single OCT detector 200, thecombining coupler 131 a preferably routes most of the sample light tothat detector. The splitting coupler routes 90% of source light, 4.5 mW,to reference fiber 113 a, while only 0.005 mW is required at thedetector. One could use a combining coupler 131 a that couples 0.1% ofthe reference light into the single OCT detector 200, but in manufactureit is difficult to control the 0.1% coupling factor. A preferredsolution is to use a 99%/1% split ratio in combining coupler 131 a, andtake advantage of the additional degree of freedom in tap 121 to adjustthe reference power. Nominally, tapping 89% of the power form referencefiber 113 a will provide an appropriate reference level of 0.005 mW atOCT detector 200, in this example.

As an alternative to adjusting the tap ratio of optional tap 121, onecan adjust the reference level by including attenuating fiber (U.S. Pat.No. 5,633,974) in the reference path.

FIG. 9 illustrates one of several other possible interferometertopologies. Low coherence light from source 101 is divided by coupler111 b between fiber 112 b and reference fiber 113 b. Sample-routingcoupler 141 further splits the light from fiber 112 b between monitor142 and sample fiber 112 c. Light in sample fiber 112 c is delayed byapparatus 114 and scanned by scanner 116 across sample 30 b, and somelight scattered from the sample is returned through these devices tosample fiber 112 c. Some of the returned light, preferably a largefraction, is routed by sample-routing coupler 141 to combining coupler131 b, where it is interfered with the reference light. Again, one orboth of the two outputs of combining coupler 131 b can be used to detectthe signal for OCT.

Considering an example as for FIG. 9, with a 5 mW source, theappropriate sample power can be achieved if splitting coupler 111 bdirects 90% of the source light to fiber 112 b and 10% to the referencepath, and sample-routing coupler 141 couples 12% of the light in fiber112 b to the sample fiber 112 c. This split ratio in coupler 141 routes88% of the light returned from the sample from fiber 112 c to fiber 117b and toward the detector. The appropriate reference level is obtainedif the combining coupler 131 b couples 1% of the power in the referencefiber to the detector, allowing 99% of the light in fiber 117 b to reachdetector 200.

Preferably, the path length to the sample is changed while maintainingthe OCT beam focus and without changing the range of sample to bescanned. One practical solution in ophthalmic imaging is placement ofthe path length adjustment between the output of the interferometer andthe scanner. However, adjusting the path length can cause the OCT beamto move transversely, offsetting it from the center of the entranceaperture to the scanner. In typical scanners, this offset causes a phaseshift in the OCT beam as the beam is scanned. Such phase shifts causesignal loss or positioning artifacts in frequency-domain techniques ofOCT.

FIG. 10 a illustrates the delay stage. Light from sample fiber 112 iscollimated by lens 120 to form OCT beam 128. Alignment mirror 123 is oneway to align the direction OCT beam 128 to be parallel to the travel ofthe moveable corner cube 124; the importance of this alignment isdiscussed later. Alternatively, beam 128 can be aligned by mechanicallymoving lens 120 in conjunction with the light emitting end of samplefiber 112. A moveable corner cube is one way to vary the optical pathlength of the OCT beam, in order to approximately match the opticalgroup delays between the sample path and reference path.

Adjustment minor 125 directs the OCT beam to scanner 116. The scanner116 can be implemented using a pair of rotatable mirrors 126 and 127,which in conjunction with scan lens 155 scan the OCT beam across sample30.

If the OCT beam is not centered on the axis of rotation of scan minors126 and 127, then as these mirrors rotate the optical path length to thesample is changed, as explained for example by Podoleanu (Podoleanu, A.G., G. M. Dobre, et al. “En-face coherence imaging using galvanometerscanner modulation.” Optics Letters 23(3): 147-149 (1998)). The effectof the scanner on the sample path length is doubled because the returnpath of light scattered from the sample back to fiber 112 is alsoaffected. This change in optical path length causes a phase shift in theinterferogram. A continuous phase shift corresponds to a shift inoptical frequency, and such a frequency shift due to relative motion isgenerally termed a Doppler shift. This Doppler shift has undesirableeffects on the data collection by frequency-domain OCT techniques, asexplained by Yun et al. (Yun, S. H., G. J. Tearney, et al. “Motionartifacts in optical coherence tomography with frequency-domainranging.” Optics Express 12(13): 2977-2998 (2004)).

Adjusting mirror 125 can be tipped and tilted to center the OCT beam onthe axis of rotation of mirrors 126 and 127. The Doppler shift due toscanning can be easily be measured by the OCT system, so as to provide aDoppler signal to be nulled by adjustment of minor 125. One way tomeasure this signal is to provide a non-moving sample 30, repeatedlyscan the OCT beam across the sample, and record closely-spaced OCTinterferograms. Pairs of neighboring interferograms should be recordedfrom locations of tissue that are close compared with the opticalresolution of the scanner, so the sampled regions significantly overlap.Pairs of neighboring interferograms differ largely in the phase shiftcaused by the optical path length change associated with transversescanning. The phase shift between neighboring interferograms is thus ameasure of the phase shift associated with the scanner, and provides asignal which is zero when the OCT beam is properly centered. Scanningthe beam in alternate directions produces an alternating phase shiftassociated with the scanner, allowing one to distinguish this phaseshift from other effects, such as the Doppler shift due to unintentionalmotion of the sample. Scanning each of minor 127 and 126 separatelyproduces a phase shift proportional to the misalignment of the OCT beamoff the respective axes of rotations of these mirrors.

If the center of the beam is mis-positioned by as little as 0.5 mm, thenthe phase shift induced by rotation of the galvo is significant. Thegalvo rotates 0.7 degrees mechanical per millisecond during a 20-degreecube. The motion of the mirror at the beam center is 6 mm/s, moving 240nm during a 40 μs exposure. This motion is sufficient to causesignificant fringe washout. The misalignment tolerance follows from theacceptable sensitivity loss due to fringe washout. The sensitivity lossdue to axial motion can be found from Yun et al [Optics Express 12(13):2977-2998 (2004)] and in terms of decibels the loss in sensitivity is

2.9 dB (q Δz)²=18 dB (Δz/λ)².

Requiring the axial sensitivity loss to be less than 0.5 dB yields thatthe Δz due to minor misalignment should be less than 0.167*λ=0.14 μm forλ=f nm. During the exposure of one A-scan from in a 128×128 cubecovering 20°, we move the beam 0.16° in the patient's field of view. Fora typical optical setup, the pupil will be imaged on the scanningmirrors, but with magnification typically 2.5, so that the minors needrotate only ⅕ of the angular sweep of the beam at the patient's pupil.Thus, the scan minor rotates by 0.03° during the exposure of one A-scan.The resulting misalignment tolerance is then

0.14 μm/[2)tan(0.03°)]≈135 μm.

Note that the tolerance to lateral misalignment, between the OCT beamand the rotation axes of the scanners, scales with the size of the imageof the pupil on the scanner.

In some applications, there is a fast scan direction and a slow scandirection. For example, minor 127 may scan rows across sample 320 andmirror 126 may move less often to move the OCT beam between scan rows.In these situations one degree of freedom is relatively more importantin the adjustment of minor 125. In general there is one direction ofscan that is relatively faster than another, and in general there is onedirection for which stable alignment is relatively more difficult. Thedesign will preferably choose the more stable alignment direction to bethe direction associated with phase shifts due to the faster directionof scan.

The measured phase shifts associated with scanning each mirror 126 and127 can provide feedback to drive adjusting minor 125 to the positionthat gives a null phase shift. Such a feedback system would allow theapparatus to self align during operation, if the subjects are relativelystill.

Alternatively to feedback using OCT, the correct position of the OCTbeam can be marked by other means. For example, a beam splitter candirect a small fraction of light from OCT beam 128 to aposition-sensitive detector that is preferably close to the locationconjugate to the rotation axes of minors 126 and 127. The properposition of the OCT beam is associated with the signal values outputfrom this position-sensitive detector during a condition of correctadjustment. In operation, the system can adjust mirror 125 to restorethe signal from the position-sensitive detectors that corresponds tocorrect adjustment of the OCT beam position.

The correct adjustment of the OCT beam on the scanner can be adverselyaffected by motion of the delay stage 114. If the OCT beam 128 is notaimed to be parallel to the direction of motion of corner cube 124, thenthe transverse position of the retro-reflected beam will change upontranslation of the corner cube. FIG. 10 a illustrates two positions ofthe corner cube 124 and the respective positions of the retro-reflectedOCT beam. In FIG. 10 a, the drift in adjustment has been corrected usingthe feedback mechanisms discussed previously so that the beam positionout of the corner cube after translation of the corner cube coincideswith the position of the beam out of the corner cube before translation.FIG. 10 b illustrates two positions of the corner cube 124 and therespective positions of the retro-reflected OCT beam. Beam 133 is theretro-reflected OCT beam directed to scanning mirrors 126 and 127 beforetranslation of corner cube 124, while beam 135 is the retro-reflectedOCT beam directed to scanning minors 126 and 127 after translation ofcorner cube 124. It is illustrated here that the beam out of the cornercube is not only delayed, but it is also translated unless the entrancebeam is properly aligned.

Feedback correction of the adjustment minor 125 will be easier, andpossibly un-necessary, if the OCT beam is well aligned to the directionof travel of delay stage 114. Such alignment can be implemented byappropriate tip and tilt of alignment mirror 123. One method foralignment of minor 123 is an extension of the method used to adjustmirror 125. The phase shift associated with scanning (or the position ofthe OCT beam on a position-sensitive detector) can be measured for twolocations of corner cube 124. Minor 123 is aligned null any change inphase shift (or position-sensor signal) with motion of the corner cube124.

The delay rail that moves cube 124 is preferably mounted in aneffectively kinematic way, to avoid misalignments of the OCT beam causedby strains in the optical mounts, such as those caused by thermalexpansion. For example, FIG. 11 shows such a rail 171 mounted to plate172 via bolts 175 and 176. If the bolt holes in plate 172 becomeimproperly spaced due to thermal expansion of plate 172, then rail 171could become bent out of plane of the figure; and the OCT beam could beshifted transversely from its desired location. Cutting holes 191 and192 into plate 172 allows the remaining plate material in 193 to flex,so that the bolts 175 and 176 can maintain the proper spacing to matchtheir holes in rail 171, and so that rail 171 remains straight. Themounting is effectively kinematic because it effectively relaxes theaxial (z-axis in the figure) constraint on the position of rail 171.Without holes 191 and 192, bolts 175 and 176 imposed competingconstraints on the axial position of the rail; the flexibility ofmaterial 193 relieves the redundant constraint.

Alternatives to expansion holes include: applying a heat sink to thesupport plate 172 or manufacturing the system so that excess heat doesnot accumulate at plate 172. Alternatively, plate 172 can bemanufactured from materials with sufficient strength to support therail, but a low enough expansion coefficient to prevent unacceptableflexing of the rail 171. Alternatively, combinations of these mechanismsor others can be used to ensure proper alignment of the corner cube 124during system operation.

Corner cube 124 is often constructed from solid glass, using internalreflections to guide the beam. The remaining surface of the corner cubecan produce weaker reflections. Such reflections are undesirable in anOCT system because if they either return to the fiber 112, or followpaths parallel to the main OCT beam, they can produce additionalinterference signals corresponding to different optical delays from thatof the main beam. The additional interference signals can result inghost images. If a corner cube is used in the longitudinal delay device,intentional misalignment or anti-reflection coating can be used toreduce reflections.

The OCT interferometer of FIG. 8 has its reference path entirely infiber, while the sample path contains some air. (Air space is requiredfor example in delay line, scan optics and working distance from opticsto the sample). The different chromatic dispersions of fiber and aircause the relative optical delay between reference path and sample pathto vary across optical frequency (see, for example, co-pending U.S.patent Ser. No. 11/334,964, filed Jan. 19, 2006, publication2006/0171503, incorporated herein by reference). This variation inoptical delay, if not corrected, leads to an uncertainty in the opticalpath length to each scattering center in the sample, worsening the axialresolution of the resulting tomograms. At certain wavelengths, such asnear 1300 nm, the chromatic dispersion of fiber is near zero, so aninterferometer configuration in FIG. 1 requires no correction formismatched dispersion. For many applications of OCT differentwavelengths are preferred, such as retinal OCT in which absorption bywater in the eye would absorbs 95% of 1300 nm scattered from the retina,before that light exits through the front of the eye.

In order to manage the mismatch in chromatic dispersion, some elementsin the sample path, which tends to have lower dispersion than the allfiber reference path, can be constructed using highly-dispersiveglasses. For example, flint glass has significantly greater chromaticdispersion than optical fiber, so constructing corner cube 124 fromflint glass significantly reduces the mismatch in chromatic dispersion.Each 1 mm of flint glass substituting for crown glass in the sample pathapproximately balances the chromatic dispersion mismatch resulting formthe inclusion of 6 mm air in the sample path. Substituting sufficientflint glass for crown glass can also overcompensate, and reverse thesign of dispersion mismatch, if desired.

Previous OCT devices required balanced chromatic dispersion betweensample and reference paths. If the reference path is entirely, or nearlyentirely, in fiber and some of the sample path is in air, there istypically a mismatch in chromatic dispersion. Such devices perform bestwhen using wavelengths for which the chromatic dispersion of the opticalfiber is nearly zero. This restriction limits the device applications tothose where the operation wavelength is chosen based on chromaticdispersion properties and not based on subject penetration or imageoptimization. Optical devices can be built to compensate for dispersionbut often at the cost of optical loss, so these devices in the samplepath would typically reduce sensitivity. Alternatively, one cannumerically compensate the chromatic dispersion mismatch. Numericalcompensation has benefits as described in the above cited U.S. PatentPublication No. 2006/0171503. These benefits work best when the physicaldispersion mismatch is within bounds, so, so even with numericalcompensation some method of controlling the dispersion is desired.

Having the reference path entirely or primarily in fiber does increasethe opportunity for polarization mode dispersion (PMD) in the fiber(Raja) which causes an undesirably variability in optical path lengthwith respect to the polarization state. When building any fiberinterferometer one often has to make splices, which can fail, and tomake the lengths of the fibers correct to match the optical path lengthof the reference and sample arms. Therefore one wants to be able tore-cut and re-splice the fiber. This is typically facilitated by placingextra loops in each of the sample and reference arms, with one loop fromeach arm removed each time the fiber is re-cut and re-splice. Therefore,one wants to have a considerable number of fiber loops. This leads toadditional length of fiber and the potential for considerable PMD. Thedesire to fit the fibers in a small space increases the polarizationmode dispersion, because bending induced polarization mode dispersionincreases with smaller bend radius.

PMD can be reduced by careful routing of the fiber; for example, the PMDcaused by bends in a horizontal plane can be compensated by followingthe horizontal bend with a vertical bend that provides approximately theopposite PMD. Such local compensation has the advantage that the netbirefringence change is zero when a loop is removed, as for re-splicingof the fiber. Another advantage is that the compensating birefringencehas the same temperature-dependence as the birefringence to becompensated, as they arise from the same physical cause. FIG. 12illustrates one way to achieve these compensating bends 144. Each passof fiber through the configuration drawn approximately compensates thebirefringence in a larger fiber loop (not shown) having a total bend of360-degrees with bend sections having 24 mm bend radius. The 201-degreebends of radius 14 mm fiber-bend in an orthogonal plane haveapproximately equal-magnitude birefringence of opposite sign.

In summary, the interferometer configurations of FIGS. 8-12 and theirequivalents, make efficient use of light returned from the eye, comparedto a Michelson interferometer. Such configurations enable settingappropriate reference signal levels at the detector. By using a variabledelay in the sample arm, no reference arm reflector is required, whichreduces the number of fiber-couplings and avoids delay-dependentvariations in the reference arm signal level. No fiber moves or bendswith this configuration. Although various embodiments that incorporatethe teachings of the present invention have been shown and described indetail herein, those skilled in the art can readily devise many othervaried embodiments that still incorporate these teachings.

Method of Patient Alignment for Multifunctional Fundus Imaging

The following embodiment included in one variation of the presentinvention describes a method of patient alignment for fundus imaging.This embodiment uses a suitable headrest, like the one described in U.S.patent application Ser. No. 10/843,767, filed May 12, 2004 (publication2005/0254009) which is incorporated herein by reference. In this method,whose optical paths are shown in FIG. 13, one moves the patient headrelative to an ocular lens 162 to set the human pupil at the entrancepupil of the instrument, then moves the ocular lens and patient headtogether to correct for refractive error. In FIG. 13, the vertical line165 left of the eye and ocular lens indicates the position of theretinal image formed by the ocular lens; this line is the retinalconjugate. The figure illustrates a system with a large number ofoptical components to the left of a retinal conjugate, with the ocularlens and eye to the right of the retinal conjugate. In the descriptionbelow, the terms headrest and chinrest are used interchangeably,referring to a suitable headrest as described in U.S. Patent Publication2005/0254009 capable of head support and functionality that moves thepatient's eye 30 at least along the optical axis, denoted as the z axis.

The chief ray of the scanning OCT beam and the rays of light used in thefundus imager both form ray pencils with a vertex at the center of theentrance pupil of the instrument. The scanning galvanometers in an OCTscanner or scanning ophthalmoscope determine the location of the vertexof the set of chief rays in the scanning beam. To get the beams into theeye the entrance pupil of the instrument must overlap the pupil of theeye.

It is advantageous to simultaneously maintain a focused image of thepupil of the eye for guidance in positioning of the eye so that the OCTand fundus microscope optical paths pass through the pupil of the eye.

The refractive error of the human eye varies over a range ofapproximately ±20 diopters. Therefore, there is a need to focus any OCTsample beam and the imaging optics of a fundus microscope to compensatefor the refractive error of the human eye.

While making these two adjustments, it is advantageous to keep theworking distance small (for better field of view without excessive sizeof optics) but yet large enough for patient safety. This leads todesigns where the entrance pupil is at a fixed, safe, distance from theclosest lens to the patient.

The Visucam non-mydriatic fundus camera, uses a separate off-axis irisview for alignment. The two adjustments are 1) camera-to-patientdistance to set the working distance, guided by the iris camera and 2)refractive correction by moving a lens within the camera.

Use of an ocular lens with a slit-lamp comprises moving thebiomicroscope portion of the slit lamp to focus on the retinal imageformed by the hand-held ocular lens.

Fundus cameras typically move an internal lens for compensation ofrefractive error. Typically a retinal conjugate is formed in theinstrument, at a location depending on the patent's refractive error; atthis location the pupil of the eye is typically imaged at infinity. Amoveable lens within the instrument is moved to focus on this retinalconjugate. The pupil of the eye is typically imaged at the back focalplane of this imaging lens.

U.S. Pat. No. 5,537,162 describes how to move the beam scanningmechanism with the moveable lens, so as to keep the vertex of bundle ofchief rays of the scanning beam at the back focal plane of the movinglens.

Rather than move the moveable lens and beam scanner, we move the ocularlens and the patient together so that the retinal conjugate is formed ata standard location with respect to the remainder of the optics of theinstrument.

An alternative solution would be to use a variable-power 1:1 relaysystem to re-form the retinal conjugate at a standard location. Anotheralternative is to use moveable minors to fold the optical path (in theshape of a trombone, for example) and extend the optical distance usingthe moveable minors so as to bring the retinal conjugate to a standardlocation.

The configurations described here have the advantage that the angularmagnification, from the human pupil to the scanning mirrors, remainsnearly constant in the face of compensation for refractive error. Thisfeature means allows the scan range of the OCT beam, in terms of anglein the visual field, to be determined based on the turning anglesscanning mirrors, without need for correction based on the motion oflenses for refractive error compensation.

FIGS. 14 a and 14 b provide a table listing the alignment steps forimaging, 14 b providing optional steps. Experience has shown that allthese steps can be conducted pretty much in any order, with theexception that the working distance has to be set first. Otherwise, theadjustments can be performed in any order.

All motor positions (chinrest x, y, z, ocular, polarization, z-motor)can be recorded for every patient and restored upon repeat visits.

The preferred optical coherence tomography device will contain a featurecalled “pupil following”. This is not pupil tracking, but rather amechanism that moves the head (and therefore the pupil) when thefixation target is moved.

When we move the fixation target in the optical coherence tomographydevice, the patient rotates their eye to follow the fixation target.While they do so, their pupil shifts because the center of rotation isbehind the pupil. Therefore the chinrest has to be moved sideways inorder to compensate for this.

The current implementation does the following:

-   -   1. It moves the fixation target continuously (rather than in one        large step where the fixation target suddenly is located        somewhere else and the patient has to search)    -   2. It compensates pupil shift based on a simplified eye model        and rotation.        For a “nominal” patient's eye the operator would never see the        pupil move at all. For a real patient there is some adjustment        necessary, but it helps.

Additionally, a corneal scan can be performed. One can insert a flip-indiverging lens 180 (FIG. 1) after the galvanometers, so as to form avirtual point source near the pupil conjugate. This results in a beamwaist near the pupil of the subject. One can set the power of the lensso that the beam waist is on the cornea of a typical eye, and move thez-motor by a typical eye length simultaneously with addition of thelens, so as to quickly switch between retinal and corneal OCT imaging.

When using this method, the iris view and LSO image are not disturbed,and the patient continues to see the fixation target.

Method for Combining B-Scans (“Thick B-Scan”)

SD-OCT greatly enhances data acquisition speed by simultaneouslyacquiring position and scattering intensities for all scatterers alongan A-Scan. The time savings may be used to simply allow fastercompletion of the same exams or the time may be used to acquire moredata, such as acquisition of higher density volumes. Acquisition of moredata provides the opportunity to combine data in order to improve somefeature or parameter, such as combining data by spatially compounding inorder to reduce speckle. OCT-tomograms generally suffer from degradedimage clarity due to image speckle and noise. Structures whosedimensions approach the resolution limit of the imaging system displayspeckle discontinuities. For example, the external limiting membrane inretina cross sections shows speckle discontinuities when imaged usingmedium resolution OCT. Speckle reduction is generally achieved bycompounding the image cell using various data acquired by means thatvary the speckle property, generally either frequency compounding (byviewing the speckle generating cell by means of a different opticalfrequency) or spatial compounding (by viewing the speckle generatingcell from a different spatial location, usually a different angle.) (SeeU.S. Pat. No. 6,847,449 and Schmitt, J. M., S. H. Xiang, et al. “Specklein Optical Coherence Tomography.” Journal of Biomedical Optics4(1):95-105 (1999).) This embodiment describes a method of specklereductions which combines elements of adjacent A-scans to produce aspeckle reduced B-scan with reduced noise and enhanced visualimpression. A B-scan resolution cell is a cell within the B-scan that isresolvable in the displayed image.

While this embodiment generally derives one or more A-scans from acollection of A-scans, its implementation and speckle reductionadvantage is more easily described as a method of determining a B-scanfrom one or more B-scans. In its simplest instantiation, a single B-scanis created by over-sampling the image region. (While over-sampling theoriginal B-scan is not necessary, it is easy to visualize compoundingwithout resolution reduction using over-sampled data.) A new B-scan isderived from the image data of the acquired B-scan by laterallyfiltering the acquired B-scan at each depth. The new B-scan may bedecimated either after or during the filtering step so that it is nolonger over-sampled. The resulting B-scan is speckle reduced along thelaterally filtered direction, but retains specular features acquired inthe transverse direction (orthogonal to the B-scan). For transversesmoothing in the direction orthogonal to the B-scan (and the creation ofa “thick B-Scan”), one or more adjacent B-scans can be used. Data at afixed depth in an A-scan can be combined with data from the same depthin other A-scans. Algorithmically, these combinations are simpler whenthe B-scans are acquired in a fixed grid of parallel planes (B-scans),as in FIG. 15 a. However, when combining data for speckle reduction, theresulting image contains fewer, or at least different, artifacts whenthe A-scans are not rigidly oriented on a regular grid. FIGS. 15 a-d areused to illustrate compounding B-scans. In FIGS. 15 a-d, the dotsrepresent A-scans and lines represent B-scans. FIG. 15 a shows thetraditional scan pattern. In FIG. 15 b, adjacent B-scans are acquiredwith the A-scan acquisition shifted by 50% of the spacing betweenA-scans in one of the B-scans. The shift of 50% of the spacing betweenA-scans is particularly advantageous when precisely two B-scans are usedto acquire a new, interpolated B-scan. FIG. 15 c represents combiningthree (3) rows to form a single B-scan while FIG. 15 d shows three (3)columns being combined to form a single B-scan.

In general, A-scans from M rows and N columns can be combined to form asingle computed A-scan. One method of combining is bi-linearinterpolation. An alternative combination is obtained if a median filteris used. Alternatively, the value of any depth point in the computedA-scan can be viewed as a weighted sum of the neighboring A-scans. Theweighted sum can include depth points. In general, the weights should beset so that the majority of the support for each computed pixel lieswithin one or two speckle diameters along each axis. The smaller thescope of this support, the greater the resolution (though this techniquecannot improve the resolution beyond that of the imaging system), whilethe larger the scope of this support, the greater the speckle reduction.Two A-scans are speckle diverse if they are separated by approximatelymore than ½ the diameter of a speckle cell. Preferably speckle diverseA-scans are separated by a speckle diameter, however, smallerseparations can achieve some speckle reduction. Similarly, a collectionof A-scans are speckle diverse at a point in a direction if thecollection contains A-scans which are separated by approximately morethan ½ the diameter of a speckle cell in that direction. Again,preferably they are separated by a speckle diameter.

The combination can be performed either during acquisition or postacquisition. FIGS. 16 a and 16 c represent patterns expressly designedfor combination during acquisition. FIG. 16 b is included to depict thedifference in the data collected using the prior art acquisitionsequence 16 b and the embodiment of the invention depicted in FIG. 16 c.Here the A-scan locations wiggle (are dithered) about a centerline. Thespan of the wiggle is greater than a speckle diameter, as shown in FIG.16 c. As depicted, any 4 A-scans are speckle diverse both tangent to andorthogonal to any plane orthogonal to the y-axis. Here the compoundingcan be performed during acquisition with only a required memory of a fewA-scans. Combination can limit computation by using a simple FIR filter,such as a boxcar or lowpass filter, or non-linear filter, such as amedian filter. Combinations may also use be complex, using higher orderstatistics derived from neighboring A-scans. The primary advantage ofcompounding during acquisition is that the data is acquired atsufficiently high speed that motion during acquisition can be ignored.For the target SD-OCT system described herein, typical modulationparameters are nominally on the order of 10 μm for the period of thewiggle and 10 μm amplitude. A range of 5-20 μm is typical. Clearly, anydifferences in the system design affecting the resultant speckle sizewould also affect the nominal modulation parameters.

The deliberate decrease of discrimination in the orthogonal direction toa B-scan may constitute a new OCT display modality of tomographic data(“thick B-scan”). The thick B-scan modality is useful for viewinglayer-like structures with a thickness close to the speckle limit.Further new displays combine the thick B-scan and one or more standardB-scans. For example, a thick B-scan derived from data from three (3)B-scan planes, say B1, B2, and B3, can be displayed essentially alsoshowing B₁, B₂, and B₃. We display the intensity of the thick B-scanwith the hue determined by B₁, B₂, and B₃; where the hue is blue if theintensity of B₁ is closest to that of the thick B-scan, the hue isyellow if the intensity of B₂ is closest to that of the thick B-scan,and the hue is red if the intensity of B₃ is closest to that of thethick B-scan. Any such display provides a suitable presentation of thedeviation of the contributing B-scans from the combined one and can evenoffer some spatial interpretation of the two dimensional data even in aprinted version.

This embodiment enables the selection of a B-scan with: reduced speckleand noise; no significant loss in lateral resolution in the direction ofthe B-scan; no increase in the number of detectors or, in some cases,scan-time; optionally increased density of A-scans in the presentedB-scan; and optionally adjustable lateral resolution orthogonal to theB-scan direction. While B-scans are nominally thought of as planarsections, they can be any curved surface. The most typical B-scans areplanar cross-sections and circle scans. Circle scans are scans coveringa cylindrical surface whose perpendicular cross sections are nominallycircular. Circle scans are particularly useful for determining retinalnerve fiber layer health, where the macular nerve head lies within thecircle scan (nominally at the center of the cylindrical surface of theB-scan.)

Adaptive Compensation of Galvo Response

The motors of optical scanning galvanometers (“galvos”) mechanicallywear out after being scanned back and forth many times (up to billionsof cycles). Before they fail completely and catastrophically, theirbearings and lubricants may deteriorate gradually over a long period oftime. These motors are typically driven by servo amplifiers that attemptto minimize the difference between the actual galvo motor position and acommanded position, provided as either an analog or a digital signal. Inophthalmic scanning, image quality and repeatability is directlydependent on galvo performance.

Two embodiments which ensure consistent scanning performance (i.e.consistent position response for a given command sequence) over thelifetime of an application are:

-   -   1) Characterize the closed-loop response and adapt the loop        filter of the servo in order to keep the closed-loop response        constant, and    -   2) Characterize the closed-loop response of the galvo system and        adapt the command signal given in order to achieve the desired        position response.

In both of these embodiments, it is essential to characterize thecomplete closed-loop response of the system. This can be done, forexample, by having some software on an instrument that gives a whitenoise command to each galvo and digitizes the position response. TheFourier transform of the position response would provide the closed-loopfrequency response.

In approach 1), a mechanism is needed to adjust the tuning of the servofilter. If the servo is digital, this can be accomplished using softwarerunning on an instrument. This can also be accomplished if the servo isanalog but has digitally settable potentiometers in the servo circuit.Appropriate adaptive filter algorithms can be achieved using knowntechniques.

In approach 2), the closed-loop response is allowed to change over time.The desired command signal at any given time can be determined byapplying the inverse of the closed-loop response to the desired positionresponse.

In practice, it is best to utilize high-acceleration, asymmetric galvowaveforms for the purpose of “clearing” the galvo motors. Thesewaveforms are designed to cause the balls in the galvo bearings to skidto a new position and avoid pit formation where the balls rock back andforth in the galvo raceways. Reliability testing has shown that theseclearing moves can prolong the life and performance of the galvos.

Calibration Test Eye

In order to align, calibrate and test an ophthalmic instrument, it isdesirable to have an artificial test eye. Various artificial eyes havebeen used throughout the ophthalmic industry, some very simple with poorimaging quality, others are more complicated, imitating the structure ofthe human eye (cornea and lens) and achieving high optical quality atgreat cost.

FIG. 17 illustrates a test eye using one single refractive surface and astop. This arrangement achieves very high optical quality over a47-degree full field of view.

One embodiment is obtained using a single piece of optical glass, with astop placed in the center of curvature of the first surface to avoidcoma, astigmatism and lateral color. The imaging surface is curved tomatch field curvature and any pattern on the imaging surface isgraduated in arc-mm to compensate for the distortion.

An alternate embodiment is obtained using aspheric surfaces to furtherreduce spherical aberration.

Co-Focus of Fundus Imager and Fixation Target

Ophthalmic instruments imaging the retina (fundus camera, LSLO, CSLO,OCT) use an internal fixation target to align the eye. It is desirableto co-focus the imaging optical path with the internal test targetoptical path so that the fundus image seen by the optician is in focusat the same time as the fixation target seen by the patient. Thefollowing paragraphs describe a method and apparatus which achieves thisresult.

Standard practice is to focus the imaging path and the fixation path inthe same plane. FIG. 18 illustrates the preferred focus describedherein. In this embodiment, the preferred focus of the fundus imager isanterior to the photoreceptor layer of the eye. The preferred focus forthe fundus imager is at the blood vessel layer, while the preferredfocus for the fixation target is at the photoreceptor layer. Therefore,we focus the imaging path and the fixation path in different planesseparated by a distance. This enables the patient to see a sharperfixation target during the eye examination when the fundus is beingimaged by the doctor. The sharper fundus image improves the patient'sattention on the fixation target, thereby decreasing eye motion andcreating a higher quality fundus image by reducing motion artifacts.

The standard design of camera lenses achieves near zero longitudinalaberration across all wavelengths within its design parameters. Oneembodiment achieves the desired result by implementing a camera lenswith a known positive longitudinal chromatic aberration. That is, thelens longitudinal aberration is the sum of human eye chromaticaberration and the desired focus shift at the specified wavelength.

FIGS. 19 a-c compare an ophthalmic instrument aligned with the prior artapproach to an ophthalmic instrument aligned using the disclosed method.FIG. 19 a shows the actual alignment of the Prior Art system with afundus imaging system and a fixation target which is not chromaticallyaligned. The vertical arrow of FIGS. 19 a-c represents the refractivepower of the cornea and crystalline lens, combined as a single lens.While the system is ostensibly designed to focus both the fundus imagerand the fixation target systems in the same plane, because the chromaticaberration of the eye optical components is not accounted for, theactual focus has the infrared fundus imager focus posterior to thevisible fixation target. That is, the focus of the fundus imaging systemis too deep. FIG. 19 b shows the same system with the fundus imagingsystem properly chromatically aligned, but without the preferred fundusimaging focus alignment. FIG. 19 c shows the same system with thepreferred fundus imaging focus alignment. The desired alignment can beaccomplished during manufacturing either by using different lenses foraligning different optical systems or a single lens with the desiredoptical properties of each system. The first alternative has theadvantage of utilizing readily available lenses and the disadvantage ofchanging lenses during final system alignment. The latter alternativehas the advantage of not requiring manufacturing to change lenses in thetest fixture during alignment with the disadvantage of requiring aspecial purpose lens which may require re-design in case of systemcomponent changes. FIG. 19 c provides the design parameters for one suchspecial purpose lens

Other imaging systems combined in one instrument are commonly aligned ina common plane in the prior art and can also benefit from furtherembodiments of this invention. For instance, in color fundus cameraswith imaging arrays, multiple sensors are used for different wavelengthsof illumination. In those cases, it is possible to adjust the axialpositions of the sensors relative to each other, so that each sensor isoptically conjugate to the source of scattered light. If light from allwavelengths is scattered from the same depths, then the optical systemis compensating for chromatic aberration to have all wavelengths at thebest focus simultaneously. Alternatively, one sensor configured toreceive visible light may be conjugate to layers anterior to the retinalpigment epithelium, and a sensor configured to receive near-infraredlight may be conjugate to the choroidal blood vessels posterior to theretinal pigment epithelium. Therefore, different layers of the retinamay be imaged on different sensors simultaneously.

In the case of retinal OCT systems, a fundus camera and confocalscanning optics provide simultaneous imaging of the retina. For example,the Stratus OCT (Carl Zeiss Meditec, Inc., Dublin, Calif.) employs afundus viewing system that is color-corrected so that both the OCT beamand fundus viewer, for a range of visible and near-infrared wavelengths,are both conjugate to the same depth in the retina.

Scanning imagers, such OCT scanners, may not provide sufficient speed toproduce real-time images to allow technicians to align the OCT scan areawith the desired region to be imaged, for instance the foveal region ofthe retina. To assist this placement, a continuously displayed image ofthe retina is desired. Illumination with near-infrared light, forinstance in the range of 700-900 nm, provides an image of the funduswithout causing patient discomfort and/or constriction of the pupil. Thescattering efficiency and absorption in the retinal layers above the RPEis relatively low in the near infra-red. Therefore, it is difficult toproduce images of retina in these layers with near infra-red light.However, these layers are clinically very important, for instance tocharacterize retinal pathologies such as macular holes, and scanningsystems, such as OCT, are often focused on these layers.

One option is to provide an infra-red fundus viewer, such as with anarray sensor in a configuration similar to that in Stratus OCT. Theinfra-red image in this case has the best contrast below the RPE, wherethe choroidal vessels are imaged. Therefore, the fundus viewer isadjusted so that it is conjugate to the choroidal vessels when the OCTscanner is conjugate either to the RPE or to layers anterior to the RPEsuch as the inner plexiform layer. This can be achieved, for instance,by first adjusting the OCT and fundus viewers to be conjugate to thesame layer in the retina or to a test fixture, then by shifting theaxial location of the fundus viewer sensor so that it is conjugate aspecific distance posterior to the RPE corresponding to tissue posteriorto the RPE. This distance would typically be in the range of 0.2 to 1.0mm. For example, in reference to U.S. Pat. No. 7,140,730 B, FIG. 3, theCCD could be shifted along the optical axis as described above.Alternatives to the alignment method, such as using a flip-in spacer tomake the adjustments, and alternatives to the design adjustment, such asshifting lenses instead of the sensor, are obvious to those skilled inthe art. Providing best focus at different depths simultaneously offerstwo advantages. First, the best focus of the fundus image, occurringposterior to the RPE, corresponds to the best focus of the OCT scannerat the desired depth so that the user can use cues from the fundusimage, such as the sharpness of the image, to adjust the focus of theOCT. Second, the fundus image provides the best possible features foruse as landmarks in placing OCT scans.

Another option is to provide two scanning imaging systems. The firstsystem, such as an OCT scanner, is slower than the second system, suchas a scanning laser ophthalmoscope (SLO), a line scanning laserophthalmoscope (LSLO), or a line scanning ophthalmoscope (LSO). Thesecond system provides video-rate images of the area to be scanned, suchas the foveal region of the retina. When the second system is a confocalimager, even near-infrared light can be used to provide good contrastimages of the blood vessels anterior to the RPE. To improve the contrastand sharpness of those images further, the second scanner is adjusted sothat it is conjugate to a layer somewhat anterior to conjugate of thefirst scanner. For example, the second system may be conjugate to bloodvessels anterior to the RPE when the first system is conjugate to theRPE layer. This offset is typically in the range of 0.2-0.5 mm dependingon the expected state of pathology in the eye.

Another option is to provide an infra-red fundus viewer, such as with anarray sensor, with an OCT scanner in a configuration similar to that inStratus OCT, where both the fundus viewer and OCT scanner are opticallyconjugate to the same layer. The new instrument achieves the desiredseparation of focal planes by first aligning the infra-red fundus viewerto its best focus on the choroid (tissue posterior to the RPE). When youare ready to capture OCT data, we then automatically shift the focus ofboth the infra-red fundus viewer and the scanner to the desired scandepth. This shift can be achieved, for instance, by motorizing at leastone of the two imagers and then using the motor to shift the sensor or alens axially before acquiring the scanned image. Alternatively, thisfocal shift can be accomplished by flipping in a lens or swapping outone lens for another to shift the focal plane the desired distance.

FIG. 20 shows, in pictorial form, a conventional arrangement of twoimaging subsystems. A scanning imaging subsystem includes a pencil beamlight source 1001 and scanning mechanism 1002, scanning lens 1111,chromatic beam splitter 1050, focusing lens 1150, lens of the eye 1160and a scanning area which sweeps out a surface bounded above by 1205,bounded below by 1215 and including focal point 1210. As the scanningmechanism sweeps the pencil beam across the scanning area, the focalpoint 1210 moves about a region on the retina. A second imagingsubsystem is a fundus camera with area illumination light source 1021and beam splitter 1040 (in this case depicted by a pin-hole minor). Thesecond imaging system includes, common to the scanning imagingsubsystem, optics path elements: chromatic beam splitter 1050 focusinglens 1150, lens of the eye 1160 and a surface of points near the retina,including point 1221, which will be brought to focus on detector 1030.When one says that the two imaging systems are focused at the sameplane, what is meant is that the scanning area swept by the scanningimager is essentially the same (or contains or is contained in) surfacewhich the second imaging systems brings into focus at detector 1030.Lens system 1133 functions to focus the fundus camera image on the CCDcamera detector 1030.

In one embodiment of the current invention, the CCD camera detector 1030is moved to be conjugate to a point posterior to 1210.

FIG. 21 illustrates an arrangement of two imaging subsystems. Theresidual portion of the first imaging system which is not depictedresides in 1010. This can be a beam imaging system, either scanning orfixed, or an area illuminating light source, like a fundus illuminatingsource. The residual portion of the second imaging system which is notdepicted resides in 1020. This can also be a beam imaging system, eitherscanning or fixed, or a line or area illuminating light source, like anSLO or fundus illuminating source, respectively. The first imagingsubsystem is conjugate to the focus at 1210 while the second imagingsubsystem is conjugate to the focus at 1220. The lens of the eye 1160,the lens 1150, and one side of the chromatic beam splitter 1050 are partof a common optical path for both subsystems. Lens 1110 completes thefocus so that the first imaging subsystem is conjugate to 1210. Lenses1120 and 1130 complete the focus of the second imaging subsystem so thatdetector 1030 is conjugate to 1220. Beam splitter 1040 (here depicted asa pin-hole minor) acts to redirect the source for the second imagingsubsystem to align with the common optical path, while allowingreflected light back to detector 1030.

FIG. 22 illustrates a simplification of FIG. 21. The effect of lenses1150 and 1160 are combined into lens 1150′. The lens 1120 is replaced bylens 1120′ so that the light is not collimated at the pinhole mirror.Lens 1130 is replaced with lens 1130′ so that 1220 is conjugate todetector 1030. Lens 1110 is replaced with lens 1110′ so that 1210 isconjugate to the detector of the first imaging subsystem.

FIG. 23 illustrates yet another configuration of the invention. Here thesecond imaging subsystem focuses anterior to the focus of the firstimaging subsystem in order to illustrate that either subsystem can imageanterior to the other. The lens 1150′ is replaced by lens 1150″, so that1220 is essentially conjugate to the pinhole minor. Lenses 1120″ and1130″ are consistent with lens 1150′ and are designed so that 1220 isconjugate to detector 1030. Lens 1110″ is consistent with lens 1150′ andis designed so that 1210 is conjugate to the detector of the firstimaging subsystem.

In FIGS. 21, 22, and 23, only the configuration is shown, where thefocal points 1210 and 1220 are not aligned. As previously described,this can be accomplished in various ways, including the changing out oflenses or movement of the detector.

It should be understood that the embodiments, examples and descriptionshave been chosen and described in order to illustrate the principals ofthe invention and its practical applications and not as a definition ofthe invention. Modifications and variations of the invention will beapparent to those skilled in the art. The scope of the invention isdefined by the claims, which includes known equivalents andunforeseeable equivalents at the time of filing of this application.

1. (canceled)
 2. A method of generating a cross-sectional image of asurface within a tissue using an optical coherence tomography (OCT)system comprising: acquiring a plurality of OCT A-scans from within thetissue, at least some of the A-scans being orthogonally displaced fromthe surface to be imaged; generating a cross-sectional image of thesurface within the tissue by compounding a plurality of A-scans, whereinat least a subset of the A-scans are selected from positions displacedorthogonally with respect to the surface, such that speckle in thecross-sectional image is reduced compared to a cross-sectional imagegenerated only from A-scans along the surface; and displaying or storingthe resulting cross-sectional image.
 3. A method as recited in claim 2,wherein two or more of the plurality of A-scans are separated by atleast approximately half a diameter of a speckle cell.
 4. A method asrecited in claim 2, wherein two or more of the plurality of A-scans areseparated by at least approximately a full diameter of a speckle cell.5. A method as recited in claim 2, wherein two or more of the pluralityof A-scans are separated by between approximately 5 and approximately 20microns.
 6. A method as recited in claim 2, wherein the cross-sectionalimage corresponds to a substantially planar surface in the tissue.
 7. Amethod as recited in claim 2, wherein the cross sectional imagecorresponds to a substantially cylindrical surface in the tissue.
 8. Amethod as recited in claim 2, wherein the cross sectional imagecorresponds to a curved surface in the tissue.
 9. A method as recited inclaim 2, wherein the compounding step includes bi-linear interpolationof two or more of the plurality of A-scans.
 10. A method as recited inclaim 2, wherein the compounding step includes median filtering of twoor more of the plurality of A-scans.
 11. A method as recited in claim 2,wherein the compounding step includes weighted summing of two or more ofthe plurality of A-scans.
 12. An optical coherence tomography (OCT)system, the OCT system including a light source generating a light beam,a sample arm, a reference arm and a detection arm, said OCT system forgenerating a cross sectional image of a surface within a tissue, saidOCT system comprising: optics for scanning the light beam to a pluralityof positions in an X/Y plane; a detector coupled to the detection armfor generating output signals in response to light collected from thesample arm and the reference arm and wherein an A-scan corresponds to areflectance distribution as a function of depth (Z) at each X and Yposition; and a processor for controlling the scanning optics and forreceiving the output signals generated by the detector, said processoroperating to acquire a plurality of OCT A-scans within the tissue, atleast some of the A-scans being orthogonally displaced from the surfaceto be imaged, said processor generating a cross-sectional image of asurface within tissue by compounding a plurality of A-scans such thatspeckle is reduced compared to a cross-sectional image generated onlyfrom A-scans along the surface.
 13. An apparatus as recited in claim 12,further comprising a display for displaying the generatedcross-sectional image.
 14. An apparatus as recited in claim 12, whereintwo or more of the plurality of A-scans are separated by at leastapproximately half a diameter of a speckle cell.
 15. An apparatus asrecited in claim 12, wherein two or more of the plurality of A-scans areseparated by at least approximately a full diameter of a speckle cell.16. An apparatus as recited in claim 12, wherein t two or more of thecompounded A-scans are separated by between approximately 5 andapproximately 20 microns.
 17. An apparatus as recited in claim 12,wherein the cross-sectional image corresponds to a substantially planarsurface in the tissue.
 18. An apparatus as recited in claim 12, whereinthe cross sectional image corresponds to a substantially cylindricalsurface in the tissue.
 19. An apparatus as recited in claim 12, whereinthe cross sectional image corresponds to a curved surface in the tissue.20. An apparatus as recited in claim 12, wherein the combining stepincludes bi-linear interpolation of two or more of the plurality ofA-scans.
 21. An apparatus as recited in claim 12, wherein the combiningstep includes median filtering of two or more of the plurality ofA-scans.
 22. An apparatus as recited in claim 12, wherein the combiningstep includes weighted summing of two or more of the plurality ofA-scans.
 23. A method of generating and displaying a cross-sectionalimage of a surface of a tissue from a plurality of A-scans collectedusing an optical coherence tomography device, said method comprisingcompounding a plurality of A-scans to generate a plurality of computedA-scans, wherein at least one of the A-scans being compounded in eachcomputed A-scan is displaced orthogonally from the surface being imaged;combining the computed A-scans to form a cross sectional image; anddisplaying the cross-sectional image.
 24. A method as recited in claim23, wherein the resulting cross-sectional image has reduced specklecompared to a cross sectional image generated from only A-scans alongthe surface.
 25. A method as recited in claim 23, wherein two or more ofthe plurality of A-scans are separated by at least approximately half adiameter of a speckle cell.
 26. A method as recited in claim 23, whereintwo or more of the plurality of A-scans are separated by at leastapproximately a full diameter of a speckle cell.
 27. A method as recitedin claim 23, wherein two or more of the plurality of A-scans areseparated by between approximately 5 and approximately 20 microns.
 28. Amethod as recited in claim 23, wherein the cross-sectional imagecorresponds to a substantially planar surface in the tissue.
 29. Amethod as recited in claim 23, wherein the cross sectional imagecorresponds to a substantially cylindrical surface in the tissue.
 30. Amethod as recited in claim 23, wherein the cross sectional imagecorresponds to a curved surface in the tissue.
 31. A method as recitedin claim 23, wherein the compounding step includes bi-linearinterpolation of two or more of the plurality of A-scans.
 32. A methodas recited in claim 23, wherein the compounding step includes medianfiltering of two or more of the plurality of A-scans.
 33. A method asrecited in claim 23, wherein the compounding step includes weightedsumming of two or more of the plurality of A-scans.